N/A
Treating cardiac arrhythmias with electromagnetically-generated heat is becoming widely accepted. Radio Frequency (RF) ablation is an important alternative to pharmacologic treatment, with high success rates in treating a wide range of atrial and ventricular cardiac arrhythmias [1-3]. A similar high success rate has unfortunately not been achieved in patients with ventricular tachycardia [4,5]. Failure to cure ventricular tachycardia associated with coronary artery disease has been attributed to the small lesion size produced with currently available RF ablation catheters. Increasing the power applied to heat tissue at depth often results in excessive temperature at the electrode-tissue interface without the desired enlargement of lesion size. Since desiccation of tissue causes an abrupt rise in impedance and limits energy transfer to the tissue [6], it is believed that the maximum effective depth of the ablated cardiac tissue with the RF method is approximately 0.5 cm. However, in the typical case the myocardial infarction lies much deeper in the ventricular myocardium [7]. In addition, it has been suggested that the zone of slow conduction mediating reentrant ventricular 2 tachycardia may be up to several square centimeters in size [8]. Technologies capable of safely heating such a large volume of tissue, such as those employing microwave energy, may be well suited for ablation of ventricular tachycardia. In contrast to heating by electrical resistance as observed during RF ablation, heating with microwaves is due to a propagating electromagnetic field that raises the energy of the dielectric molecules through which the field passes by both conduction and displacement currents. Thus, in microwave ablation, the initial volume of ablation is a result of direct electromagnetically-induced, dielectric, frictional heating and is unlikely to be limited by local tissue factors or electrode size. This mode of heating lends microwave ablation the potential for a greater depth and larger volume heating than RF ablation and should result in a larger lesion size [9, 10]. Previous research has shown that lesions created by microwave energy increase in size with increased applied power [11].
Microwave catheter-based antenna applicators have been used experimentally for cardiac ablation. These applicators may be grouped into two categories: the monopolar antennas [12, 13] and helical coil antennas [6, 14, 15]. The monopolar antennas are usually one-half tissue wavelength long, designed to radiate in the normal mode to generate a well-defined football-shaped heating pattern along the antenna length. The helical coil antenna applicator is also designed to radiate in the normal mode, perpendicular to the axis of the helix. The helix has been shown to exhibit improved uniformity and localization of heating along the radiating coil portion of the antenna compared to the monopole configuration.
It is desirable, however, to have an illuminating aperture that is as large as possible. The monopole and helix antenna applicators have radiating apertures limited by the diameters of their catheters, and as such must be often be repositioned to create a sufficiently large lesion [16].
A catheter-based microwave antenna cardiac ablation applicator has been developed which unlike previously-developed ablation catheters, forms a wide aperture that produces a large heating pattern. The antenna comprises a spiral antenna connected to the center conductor of a coaxial line, and which is insulated from blood and tissue by a non-conductive fluid filled balloon. The antenna can be furled inside a catheter for transluminal guiding. Once in place at the cardiac target, the balloon is inflated, and the coiled spiral antenna is ejected into the inflated balloon. The wide aperture antenna generates a ring-shaped power pattern. The heat generated from this deposited power is conducted through a volume larger than the spiral diameter, ablating diseased tissue. The resultant lesion profile is both wider and deeper than that of either conventionally-used RF catheter-based ablation electrodes, or that of other microwave applicators, and provides greater heating accuracy and controllability.
Unlike monopole antennas, which radiate normal to their axes, and conventional RF electrodes, which generate radial currents, loop antennas can radiate in either normal or axial modes. Electrically small current loops behave like magnetic dipoles, with electric field strongest in the plane of the loop and polarized circumferentially. However, once the loop circumference approaches one wavelength (in the medium surrounding the loop), the waves it radiates are strongest in the axial direction, with rotating electric field polarized perpendicular to this axis. With proper loop radius adjustment and surrounding medium specification, it is possible to tailor the radiation pattern, creating a ring of deposited power. Thermal conduction then xe2x80x9cfills inxe2x80x9d the ring, providing a hemi-oblate spheroid lesion shape.
An important aspect of the antenna applicator design is matching of the impedance from applicator to tissue. In practice, an unfurlable spiral formed from the extended center conductor of a coaxial feed line is used instead of a loop. The spiral will be introduced through blood vessels into the heart chamber in a compact, collapsed state, and then ejected from a catheter housing and allowed to reform a spiral shape. It has been determined that the overall length of the center conductor wire governs the antenna impedance. Previous experiments [17] demonstrated that a length of about one tissue wavelength provided the best match to a 50 ohm coaxial cable.
To first order, it is possible to approximately model the spiral antenna as a circular loop. This model gives a sense of the power deposition pattern and thus the heating profiles. For a one wavelength circumference loop, current on one side of the loop will be 180xc2x0 out of phase and flow in the opposite direction from that on the diametrically opposite side. Thus, these two currents will excite fields which constructively interfere along the axis of the loop and cancel outside the loop. For a loop positioned on a planar tissue surface, this modest focusing yields an enhanced electric field and hence increased power deposition at depth within the tissue. In high water content tissue, one wavelength corresponds to 4.5 cm at 915 MHz [18], which establishes a nominal loop diameter of 1.4 cm.
It has been determined that to produce the desired size lesion of slightly greater than 1 cm diameter and prevent tissue surface overheating, an inner region of low loss, low dielectric constant fluid surrounding the spiral was needed. This physiologically benign fluid, usually air, nitrogen, or a perfluorocarbon blood substitute is contained within a balloon surrounding the loop. An inflation tube is used to fill and drain the fluid. For therapeutic purposes, it is preferable to direct the radiated power directly into the cardiac tissue. Not only does this prevent heating of blood within the heart chamber, it also delivers more of the available power into the heart tissue. Directing the power flow is achieved by asymmetrically positioning the loop in an inflatable balloon, with less fluid in front of the loop than behind it. The loop antenna is specified with radius b placed eccentrically inside the balloon with diameter c, in a plane at a distance l from the center of the balloon. For the spiral within a balloon, the simple model of a single wavelength circumference loop breaks down, and a more sophisticated moment method analysis is required. Repeated experimental and numerical trials concluded that a smaller diameter spiral performs better in the balloon-enclosed environment. The best radius was found to be b=0.5 cm (about 70% of the nominal radius), with balloon radius c=1 cm.